Optical Coherence Tomography (OCT) is an interferometric technique that can provide images of samples including tissue structure on the micrometer scale in situ and in real time (Huang, D. et al., Science 254, 1178-81, 1991). OCT is based on the principle of low coherence interferometry (LCI) and determines the scattering profile of a sample along the OCT beam by detecting the interference of light reflected from a sample and a reference beam (Fercher, A. F. et al., Opt. Lett. 13, 186, 1988). Each scattering profile in the depth direction (z) is reconstructed individually into an axial scan, or A-scan. Cross-sectional images (B-scans), and by extension 3D volumes, are built up from many A-scans, with the OCT beam moved to a set of transverse (x and y) locations on the sample.
Many variants of OCT have been developed where different combinations of light sources, scanning configurations, and detection schemes are employed. In time domain OCT (TD-OCT), the pathlength between light returning from the sample and reference light is translated longitudinally in time to recover the depth information in the sample as illustrated in FIG. 1. FIG. 1 shows a prior art TD-OCT system with a moving reference mirror 103. Also shown is a depiction of the type of data that is produced: 101 showing an A-scan, and 102 showing a 2D B-scan (composed of multiple A-scans).
In frequency-domain or Fourier-domain OCT (FD-OCT), a method based on diffraction tomography (Wolf, E., Opt. Commun. 1, 153-156, 1969), the broadband interference between reflected sample light and reference light is acquired in the spectral frequency domain and a Fourier transform is used to recover the depth information (see for example Fercher, A. F. et al., Opt. Commun. 117, 43-48, 1995). The sensitivity advantage of FD-OCT over TD-OCT is well established (see for example Leitgeb, R. et al., Opt. Express 11, 889, 2003; Choma, M. et al., Opt. Express 11, 2183-9, 2003).
There are two common approaches to FD-OCT. One is spectral domain OCT (SD-OCT) where the interfering light is spectrally dispersed prior to detection and the full depth information can be recovered from a single exposure as illustrated in FIG. 2. In FIG. 2, an SD-OCT system is depicted along with the type of data that is produced, first as a spectrally dispersed interferogram 201, an then how it looks after being Fourier-transformed 202. A B-scan image 203 produced by combining several A-scans is also depicted. In an SD-OCT system as illustrated in FIG. 2, light from a broadband light source 206 is split by a coupler 210 into reference and sample arms. The reference arm light travels towards a stationary reference mirror 205 where it is back reflected and returns along the same path to coupler 210. Light in the sample arm is directed towards an eye 209 of a patient by a series of optical elements. The light can be scanned over a plurality of transverse locations on the eye using a scanner 208. The back-reflected light from the tissues of the eye 209 is then collected and interfered with the light of the reference arm at coupler 210 and the interfering light is spectrally dispersed by a grating 207 onto a detector array 204. The electrical signals from the detector are transferred to a processor where the spectral interferogram is transformed into A-scans and then combined to create B-scans of the eye. The sample and reference arms in the interferometer could consist of bulk-optics, photonic integrated circuits, fiber-optics or hybrid bulk-optic systems and could have different architectures such as Michelson, Mach-Zehnder or common-path based designs as would be known by those skilled in the art. The processing could be accomplished in the data collection instrument or remote from the instrument. Parallel processing techniques could be employed to expedite processing.
The second common type of FD-OCT is swept-source OCT (SS-OCT) where the broadband source is replaced with a frequency tunable source that is swept over a range of optical frequencies in rapid cycles and the resulting signal is detected in time using for example, a balanced detector, therefore encoding the spectral information in time. In traditional point scanning or flying spot techniques, a single point of light is scanned across the sample. These techniques have found great use in the field of ophthalmology. However, current point scanning systems for use in ophthalmology illuminate the eye with less than 10% of the maximum total power possible for eye illumination spread over a larger area. It may not be immediately possible to significantly increase the illumination power with the current point-scanning architectures since the systems already operate close to their maximum permissible exposure for a stationary beam. Parallel OCT techniques, which spread the illumination light over a larger area on the tissue may be able to overcome this challenge. Further, the typically higher acquisition speed of such parallel systems will result in comprehensively sampled volumes which are required for applying computational imaging techniques.
In parallel techniques, a series of spots (multi-beam), a line of light (line-field), or a two-dimensional field of light (partial-field and full-field) is directed to the sample. The resulting reflected light is combined with reference light and detected. Parallel techniques can be accomplished in TD-OCT, SD-OCT or SS-OCT configurations. A number of groups have reported on different parallel FD-OCT configurations (see for example Hiratsuka, H. et al., Opt. Lett. 23, 1420, 1998; Zuluaga, A. F. et al., Opt. Lett. 24, 519-521, 1999; Grajciar, B. et al., Opt. Express 13, 1131, 2005; Blazkiewicz, P. et al., Appl. Opt. 44, 7722, 2005; Pova{hacek over (z)}ay, B. et al., Opt. Express 14, 7661, 2006; Nakamura, Y. et al., Opt. Express 15, 7103, 2007; Lee, S.-W. et al., IEEE J. Sel. Topics Quantum Electron. 14, 50-55, 2008; Mujat, M. et al., Optical Coherence Tomography and Coherence Domain Optical Methods in Biomedicine XIII 7168, 71681E, 2009; Bonin, T. et al., Opt. Lett. 35, 3432-4, 2010; Wieser, W. et al., Opt. Express 18, 14685-704, 2010; Potsaid, B. et al., Opt. Express 18, 20029-48, 2010; Klein, T. et al., Biomed. Opt. Express 4, 619-34, 2013; Nankivil, D. et al., Opt. Lett. 39, 3740-3, 2014).
Non-confocal parallel OCT methods, especially full-field OCT systems can suffer from image degradation due to the increased collection of multiply scattered light. When imaging highly scattering samples, as for example the retinal pigment epithelium (RPE) in the eye, it is beneficial to suppress multiple scattered light and therefore enable imaging of deeper structures within the sample (e.g. the choroid and sclera in the case where the human eye is the sample).
Multi-beam OCT systems as well as joint aperture OCT systems have complicated system architecture with multiple interferometers. Both space-division multiplexing OCT (Zhou et al., Opt Exp 21, 19219, 2013) and interleaved OCT (Ellerbee et al., Opt Exp 21, 26542, 2013) utilized the concept of translating the long coherence length of the source into high OCT imaging speed. Different illuminated points on the sample had different optical path lengths and the interferometric signal from each point was extracted from a different imaging depth. However, one of the critical limitations of space-division multiplexing is to separate desired signal from unwanted reflections at the optical surfaces or tissue. For example, thick samples (such as signal from vitreous) may cause overlap with imaging windows of other channels and result in artifacts. In addition, both the systems use complex sample arm designs to illuminate multiple locations in the sample (by using either splitters or virtually imaged phased arrays—VIPA) that lead to around ˜10 dB losses in the sample arm. Ellerbee et al. (2013) teach illuminating multiple points on to the sample in their technique of interleaved OCT (Ellerbee et al. US Patent Publication No. 20130215431). In their technique, roughly the full spectral width of the source is divided into P sets of unique spectrally interleaved wavelength components. Each point is illuminated by roughly the full spectral width of the source. While this allows the benefits of being able to use more power to illuminate the sample safely, the mechanisms to de-multiplex the light at the sample and detection ends add significant complexity and losses in the system.
Spectral encoded endoscopy spectrally spreads the light on to a sample (see for example Yelin et al., Opt Lett 28, 2321, 2003 and Yelin et al., Opt Exp 15, 2432, 2007 and Tearney et al., Opt Lett 23, 1152, 1998 and Tearney et al., Opt Lett 27, 412, 2002). However, this approach has significantly poorer axial resolution as only a partial spectral window is used from the available full spectral width of the source for A-scan reconstruction. Gronle et al. also demonstrated a method to spectrally disperse the light on the sample and used it for height profile measurement of the sample (Gronle et al., App Opt 50, 4574, 2011). However, this approach also compromises on the axial resolution as it does not utilize the full spectral bandwidth.